Gamma cameras, also known as scintillation cameras or Anger cameras, have been used for over half a century to image gamma photon emission in a technique known as scintigraphy. They are applied in nuclear medicine imaging (also known as “molecular imaging”) to view images of the interior tissues of the human body containing a distribution of medically injected, inhaled, or ingested radiopharmaceuticals emitting gamma photons. They can also be used in security screening or astronomical imaging applications to detect and image gamma photons arising from radioactive sources or astrophysical sources of x-rays and/or gamma photons. Hal Anger developed the first gamma camera in 1957 and variations on his design are still widely used today.
A typical gamma camera comprises a large planar crystal (about 40 cm×54 cm×0.95 cm) of sodium iodide with thallium doping, labelled NaI(Tl), in a hermetically sealed container that prevents moisture and visible light from reaching the crystal. The back face has a transparent glass window or light guide. The crystal scintillates when a gamma photon strikes it and the scintillation light can escape through the window, where an array of photodetectors (such as vacuum photomultiplier tube—PMT, or multichannel plate—MCP, or silicon photomultiplier—SiPM) are coupled through optical gel. The location of the gamma photon interactions in the NaI crystal are determined by a weighting of the current or voltage responses of multiple PMTs that detect the scintillation light.
A gamma camera is typically coupled with a collimator (for example: parallel or focused hole, slit-slat, rotating slat, multiple pinhole, or coded aperture) that acts as a lens for forming a projection image. In SPECT, used in nuclear medicine, a 3D image is formed by moving the gamma camera to multiple angular positions around the patient, acquiring planar projections at each position, and then reconstructing a 3D image of the gamma emission source distribution using various techniques, such as filtered back-projection or iterative maximum likelihood estimation reconstruction. OSEM, Ordered Subset Expectation Maximization, is the currently preferred method. Commercial SPECT systems most commonly consist of two gamma cameras mounted in opposition (180 degree angle) or side-by-side (forming a 90 to 120 degree angle) on a gantry with a patient table positioned between the slowly rotating gamma cameras.
This invention addresses one or more limitations of current commercial solutions and prior art. In particular, the basic technology used in gamma cameras has not changed substantially since 1957, except that computers have increased the efficiency of the data acquisition and image reconstruction. This invention concerns in part the introduction of solid-state semiconductor gamma detectors arrayed as a gamma camera. By way of illustration, pixelated Cadmium-Zinc-Telluride (CdZnTe or CZT) detectors will be referenced, although similar considerations apply to other solid-state gamma detectors, such as Cadmium-Telluride (CdTe) and Mercuric Iodide (HgI). These pixelated solid-state detectors have many advantages over the prior art consisting of a scintillator (such as sodium iodide (NaI), cesium iodide (CsI), lanthanum bromide (LaBr), etc.) and photodetector (such as photomultiplier tube (PMT), multichannel plate (MCP), silicon photomultiplier (SiPM), avalanche photodiode (APD), etc.), but they also have the disadvantage of having a higher cost than the 60-year old prior art.
This invention also applies in part to non-pixelated detectors, such as crossed-strip or position-sensitive virtual Frisch grid designs. This invention addresses one or more ways to decrease the cost of solid-state gamma cameras, primarily by sparsely placing the CZT detector modules and then compensating with novel collimation and image reconstruction. Although we will discuss CZT gamma cameras, many of the same design principles could apply to small (less than about 400 cm2 area, but preferably less than about 25 cm2 area) modular scintillator cameras and such generalizations are intended by this invention.
By way of example, we will discuss a preferred embodiment of a CZT gamma camera 22 consisting of an array of modular pixelated CZT detectors (called D-Matrix™) that we have designed and tested. The basic module is shown in FIG. 1 and consists of four CZT gamma modules (GM) 21 in a 2×2 array comprising one aggregator module (AM) 20. Each AM 20 is 4.4 cm square and provides digital signal output including 2D location (module and pixel numbers or, preferably, x and y) and energy for each gamma photon detection event. Optionally, a 3D location including depth of interaction z in the CZT crystal can also be provided. One skilled in the art will appreciate that the AM 20 includes ASICs (application-specific integrated circuits), one or more ADCs (analog to digital converters), and potentially an FPGA (field-programmable gate arrays) as well as supporting input/output electronics. Table 1 following contains the preferred characteristics of each CZT gamma detector aggregator module 20.
TABLE 1Features of a D-Matrix ™ CZT gamma detectoraggregator module (AM), 20, as shown in FIG. 1.FeatureSpecificationAggregator Module (AM)4.4 cm × 4.4 cmdetection areaPixel pitch2 mmCZT thickness0.5 cm (or optional 1.0 cm)FootprintTileable on 4 sidesEnergy range30 keV to 400 keVOperating environmental20° C. to 35° C.temperature rangeData output (list mode)x, y, energy, (optional: z [depth ofinteraction] and various event timings)APISoftware for acquisition computer interface
By way of further example, we will discuss a gamma camera 22 composed of D-Matrix™ aggregator modules 20 arrayed in a size compatible with a general-purpose SPECT system as depicted in FIG. 2. In the illustration, nine AMs 20 are arrayed in each of 12 columns, each 39.6 cm long and 4.4 cm wide. The 12 adjacent columns comprise a camera 52.8 cm wide. Thus, there are a total of nine×12=108 AM 20, and 4×108=432 GM 21. Since each GM 21 has 121 pixels, the general purpose gamma camera 22 in FIG. 2 has 52,272 pixels, each two mm square. Some of the advantages of a CZT gamma camera compared to a scintillator gamma camera can be seen in Table 2 following.
TABLE 2Comparison of typical scintillator gamma camera to a CZT gamma camera composedof an array of D-Matrix ™ CZT detector aggregator modules.NaI & PMTsCompared Metriccamerasmodular CZT camerasDetector Crystal Thickness0.95 cm thick NaI(Tl)0.5 cm thick CZT or1.0 cm thick CZT (option)Camera Useful Field of View40.6 cm × 54 cm39.6 cm × 52.8 cm (no dead edges)Intrinsic Spatial Resolution3.3 mm FWHM2.0 mm square pixelsGaussianIntrinsic Pixels per Camera20,132 with Gaussian52,272 square, no overlap(Space-Bandwidth Product)overlapIntrinsic Energy Resolution≤9.6% FWHM @≤4.0% FWHM @ 140 keV140 keV
Table 2 shows that CZT detectors can be arrayed in a camera size similar to a typical commercial gamma camera (most are about 40 cm axial×54 cm transaxial). The five mm thickness of CZT is chosen to provide a similar stopping power to a typical 0.95 cm (⅜″) thick NaI scintillator. This thickness is adequate for the most common medical isotopes, such as Tc-99m (140 keV), Tl-201 (70 keV), Xe-133 (81 keV), Ga-67 (90 keV), and I-123 (159 keV). Increasing the CZT thickness to 1.0 cm will increase the stopping power and thus the detection efficiency for higher energy medical isotopes, such as In-111 (171 & 245 keV) and I-131 (364 keV), although at considerable increase in cost.
The intrinsic spatial resolution of the example CZT gamma camera 22 is about 1.7 times better than a typical scintillator gamma camera 22 resulting in 2.7 times more pixels in the same detector area. Moreover, the pixelated CZT gamma camera has no dead edges compared to the scintillation camera which has unusable (“dead”) edges about half the diameter of the PMTs wherein the Anger position determination is ineffective. Thus the edges of a scintillation camera must typically be masked off and not used for imaging. Furthermore, a scintillation gamma camera has a nonuniform performance wherein the central field of view CFOV performs significantly better than the peripheral (“useful”) field of view UFOV. In a CZT gamma camera the performance is uniform across the full surface, so the CFOV and UFOV performance is equal. Furthermore, the CZT pixels are square and non-overlapping (except for a small amount of charge sharing), while the scintillator pixels are Gaussian with considerable overlap between neighboring pixels. Finally, the energy resolution is typically 2.4 or more times better, resulting in a better discrimination against scattered gamma photons which blur SPECT images.
We tabulated the price that our customers paid for CZT gamma detectors mounted on substrates with connectors for attachment to ASICs and other read-out electronics and determined the price per area ($/cm2) of CZT gamma detectors since 2000. The price of CZT gamma cameras has been driven down exponentially (with a half-life of about 4 years) by improvements in CZT crystal growth and sensor fabrication. The relative prices are now only two to four times greater for CZT gamma cameras compared to NaI & PMTs. The purpose of this invention disclosure is to provide additional system design steps that can offer an additional factor of two or more reduction in the cost of CZT gamma cameras, principally by sparsely populating the cameras with CZT detectors, then compensating with collimator design and optional detector and/or collimator motion to achieve similar examination times for the improved image quality that CZT gamma cameras offer.